Sub-micron surface plasmon resonance sensor systems

ABSTRACT

A sensor for detecting the presence of a target analyte, ligand or molecule in a test fluid, comprising a light transmissive substrate on which an array of surface plasmon resonant (SPR) elements is mounted is described. A multi-channel sensor for detecting the presence of several targets with a single micro-chip sensor is described. A multi-channel sensor including collections of SPR elements which are commonly functionalized to one of several targets is also described. The detectors sense changes in the resonant response of the SPR elements indicative of binding with the targets.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a divisional application of co-pending U.S. patentapplication Ser. No. 11/611,509, filed Dec. 15, 2006, which claimspriority under 35 U.S.C. §119(e) to U.S. provisional application Ser.No. 60/750,872, filed on Dec. 16, 2005, entitled “Sub-micron CavitySurface Plasmon Sensors and Their Micro-fluidic Applications”, theentire disclosure of each is incorporated by reference herein.

GOVERNMENTAL INTEREST

This invention was made with government support under grant numberIBN-0083653 and NAG2-1619 awarded by the National Science Foundation andfrom NASA. The U.S. Government has certain rights in the invention.

BACKGROUND

A significant trend in medicine is the introduction of point of care(POC) devices for rapid, bedside diagnosis. These devices enable rapiddiagnosis by first responders or medical staff for time-criticaldiagnoses, such as for indicating whether patients are presenting withcardiac symptoms. Tests have been developed for other indications, suchas infectious diseases, drugs of abuse, cerebrovascular disease, thatare intended to circumvent the lengthy processing hours and high costsaccompanying conventional in-house laboratory assays. Current POCdevices are single use only. While this is suitable for manyapplications, there is an unmet need for continuous monitoring devices.

An initial clinical need is a device that can monitor and detect thepresence of infections in intensive care patients. Currently, manyintensive care patients develop infections that are not detectedquickly, often leading to sepsis or shock and resulting in a largemortality rate. There is a significant need for a device that cancontinuously track the concentration of specific protein markers in apatient's bloodstream that are indicative of an infection, for instance.

Devices that are capable of detecting the presence of selected chemicalsor biological substances include biosensors that interact directly witha sample molecule to provide a signal identifying the test molecule.Biosensors are often functionalized chemically to make them selective.The readout can be electrochemical, as is often the case for smallmolecules (e.g. glucose), or can utilize fluorescence or other opticaltechniques for molecules such as proteins or DNA. Typical biosensors canoften operate in a continuous reading mode or can be used multipletimes, which differs from conventional laboratory assays requiring bulkreagent handling, usually yielding only a one-time test result.

The miniaturization possibilities afforded by biosensors compared toconventional laboratory assays suggests that point of care (POC) testscould provide dramatically enhanced diagnostic capabilities. Thebenefits of POC testing include: rapid turnaround which aids therapeuticdecisions; quick dissemination of test results to patients, therebyreducing physician workload and increasing patient satisfaction; reducedpaper work and simplified sample tracking; and reduced need forspecialized technicians. POC tests administered as panels providefurther significant benefits. For example, screening for several cardiacmarkers simultaneously saves time and provides useful additional data.Screens for various types of influenza would aid diagnosis compared tomore limited tests on only single strains.

Emerging applications of biosensors include food and water testing,drugs of abuse, bio-defense and “white powder” detection, and veterinarytesting, to name a few. Some of these applications have unique needssuch as the need for ultra-fast response time in conjunction withbio-defense measures, or high sensitivity necessary in food or watertesting to detect a very low number of E. Coli colony-forming units.Typical water testing products use reagents that must be incubated inflasks for 18-24 hours or longer, changing color to indicate pathogenpresence. While these products are very effective, the lengthy, 24 hourincubation time can be problematic. When the contaminated water is in apublic drinking supply, the water may be in use for extended periodsbefore a pathogen problem is detected. A product that continuouslymonitors water quality can provide a warning within minutes of actualcontamination.

Bio-defense presents unique issues as governmental and military agenciessearch for ways to rapidly and interactively detect anthrax, botulism,malaria, Ebola virus, ricin, and other potential terrorist agents.Expensive test kits are currently used by the US Postal Service thatincorporate real-time PCR to amplify and analyze crude samples obtainedfrom air or suspicious “white powder” on packages and envelopes.

A new breed of biosensors utilizes a phenomenon arising from theinteraction of light with a metal surface. This phenomenon is called“surface plasmon resonance” and embodies a charge-density (electroncloud) oscillation that may exist at the interface of two media withdifferent dielectric constants or dielectric constants of oppositesigns. This condition is usually met at the interface between adielectric (glass) and a metal (typically gold or silver). The chargedensity wave (the electron cloud) is associated with an electromagneticwave (the incoming photons), and this coupling reaches a maxima at theinterface and decays exponentially into both media. This coupling is, ineffect, a surface bound plasma wave (SPW).

This coupling cannot be excited directly by incident optical photons ata planar metal-dielectric interface because the propagation constant ofan SPW is always higher than that of the wave propagating in thedielectric. Therefore to enhance this coupling, attenuated totalreflection (ATR), prism couplers and optical waveguides, or diffractionat the surface of diffraction gratings is used. As the excitation ofSPWs by optical photons results in resonant transfer of energy into theSPW, surface plasmon resonance (SPR) manifests itself by resonantabsorption of the energy of the optical photons. Owing to the strongconcentration of the electromagnetic field in the dielectric (an orderof magnitude higher than that in typical evanescent field sensors usingdielectric waveguides) the propagation constant of the SPW, andconsequently the SPR formation, is very sensitive to variations in theoptical properties of the dielectric adjacent to the metal layersupporting SPW, namely the refractive index of the dielectric mediawhich may be determined by optically interrogating the SPR. Thethickness of the region of sensitivity varies with the wavelength offthe applied energy, but is typically about 500 nm for wavelengths in thevisible light range. The refractive index is modified by the presence ofmaterials or impurities at the surface. This is the fundamental effectthat can be used to identify the materials or impurities with greatprecision.

Metals are materials that can provide the negative sign dielectricconstant. They have a resonant mode at which the constituent electronsresonate when excited by electromagnetic radiation having the rightwavelength. Gold, in particular, has a spectrum with a resonance atvisible wavelengths around 510 nm. In the case of the attenuated totalreflection in prism couplers, the evanescent wave is sensitive to themetal surface in contact with the media within approximately 200-400 nmof the surface, enhanced by the presence of a surface plasmon wave. Suchmaterial effectively modifies the index of refraction and thus theprecise angle of critical attenuated total reflection. Interactionsbetween a bound substrate and a sample can thus be probed, measuringsmall variations in the reflection angle at maximum SPR production.

This effect can be harnessed to study binding between molecules, such asbetween proteins, RNA and/or DNA, or between proteins and viruses orbacteria. For example, a surface functionalized with a specific antibodywill probe for only one antigen (e.g. antigen A) and discriminatespecific binding from non-specific binding. That is, antigen A will bedetected but weaker interactions between the functionalized proteinbound to the surface and another antigen, say antigen B, can bedistinguished. Typically, angular resolution of a few millidegrees isrequired to discriminate between selective and non-selective binding.Thus the detection of protein A in solution as dilute as 1 pg/ml may beachieved. In addition, the reaction kinetics of the binding between thesurface protein and antigen A can be elucidated.

Most commercial SPR instruments comprise a sample introduction device orsensor that includes a semispherical dielectric prism coated with a thinlayer (50 nm) of a noble metal such as Au or Ag. This metal coating inturn is coated with molecules that will specifically bind a targetanalyte. These commercial devices further comprise a light source on agoniometric mount, an array detector, and various collimation andfiltering optics, as depicted generally in FIG. 1.

Using a semispherical prism, the angle of incidence at thedielectric/air interface is the same as at the first air/dielectricinterface where the ray from the light source enters the prism. At theprecise incidence angle at which light couples to a non-radiativeevanescent wave (surface plasmon) in the metal film, the reflectivity ofthe film decreases roughly 90% creating an evanescent plasmon fieldwhich is localized at the metal surface away from the glass. Theevanescent wave's properties depend on the properties of the medium(e.g., biomolecules) in contact with the free metal surface of thesensor. Subtle changes in the refractive index of the medium, such asthose associated with molecular absorption onto the surface, inducedetectable changes in the surface plasmon resonance angle φ. The SPRinstrument then adjusts the detector position to find this new angle andthus measures the change in SPR angle.

These types of SPR devices have a number of inherent limitationsinvolving sensitivity, sample size, complexity, and cost. Existingcommercial instruments require large, complex, and delicate moving partsin order to optimize the incident beam and detector positions. Forinstance, the goniometric mount for the light source is relatively bigand heavy, but delicate. Moreover, the light source itself must providepolarized light. Typical sensitivity limits are on the order of 10⁻⁶refractive index units which is usually sufficient to detect targetswith a concentration of 1 pg/mm² of adsorbed molecule and a size of atleast 200 Da, but is not sensitive enough to provide useful detectionfor bio-terrorism agents in concentrations of 0.01 parts per billion asrequired by certain government standards. The typical planar sensorfootprint is in the range of a few mm² ( 1/16^(th) mm² in the BiacoreFlexichip and 2.2 mm² in the Biacore 3000) which creates a technicalconstraint on the ability to miniaturize these sensors. A larger sensorarea means that more test fluid must be provided to flow over the planarsensor. Moreover, the constraints on accuracy also require more testfluid to provide sufficient molecules or microparticles to be detected.Because of an SPR sensor's macroscopic size, arrays of sensing elementsfor multiplexed analysis require sample volumes too large for mosttechnologies used for analytical integration. All of these limitationsof conventional planar sensors reduce the throughput capability of thesensors.

Additionally, most current SPR sensors require p-polarized light (i.e.,the electric vector component is parallel to the plane of incidence) andprecise alignment of their optical parts, which are comparable incomplexity to those of a tabletop spectrometer. This results in highcost, typically on the order of several hundred thousand dollars.

DESCRIPTION OF THE FIGURES

FIG. 1 is a schematic representation of the operation of aflat-substrate SPR sensor of the prior art.

FIG. 2 is an enlarged schematic representation of an SPR sensor inaccordance with one embodiment of the present invention.

FIG. 3 is an electron-microscopic image of an SPR bead sensor fabricatedaccording to the present invention.

FIG. 4 is a schematic view of a micro-fluidic chip utilizing the SPRsensor according to the present invention.

FIG. 5 is a schematic representation of an experimental set-up forevaluating the performance of an SPR sensor according to the presentinvention.

FIGS. 6 a and 6 b are graphs of the spectral performance of the SPRsensor in the experimental set-up shown in FIG. 5.

FIGS. 7 a and 7 b are graphs of the spectral performance of the SPRsensor of the present invention under further experimental conditions.

FIG. 8 is a schematic representation of a micro-fluidic SPR sensoraccording to a further embodiment of the invention.

FIG. 9 is a diagram of the sulfo-DSP reaction with the gold layer of theSPR sensors of the present invention for functionalization of the SPRsensors.

FIGS. 10 a and 10 b are diagrams of the functionalization reactionsusing Carbodiimide coupling reagents.

FIG. 11 is a diagram of micro-fluidic components mounted on amicro-fluidic SPR sensor of the present invention.

FIG. 12 is a diagram of a micro-fluidic SPR sensor of the presentinvention with micro-fluidic filtering and pre-concentration modules.

FIG. 13 is a diagram of a micro-fluidic SPR sensor of the presentinvention with mapped functionalization for detecting multiplemolecules, ligands or analytes.

FIG. 14 is a diagram of the components of a micro-fluidics SPR sensorsystem in accordance with the present invention.

FIG. 15 is a schematic representation of a sensor according to thepresent invention that is capable of simultaneously evaluating multiplechemicals.

SUMMARY OF THE INVENTION

Evanescent-wave sensors using SPR techniques for biomolecularinteraction analysis, for instance, provide several advantages,including non-intrusiveness, real-time monitoring of the binding oftarget analytes, ligands or molecules and label-free conditions. Amechanism to increase the sensitivity of SPR sensors while reducing thesize of the sensor would be very desirable, especially in the fields ofmedical diagnostics, drug screening, biomedical research, andbioanalysis. Another desirable goal is to eliminate the often fragilemechanical and optical components that add bulk to the sensor, increaseresponse time and decrease sensitivity. In accordance with one aspect ofthe present invention, the propagating plasmon wave is replaced with astationary wave or, in other words, the sensitivity of the SPR sensor isenhanced by adding shape resonance. Such a stationary wave will travelacross the active surface a number of times proportional to the qualityfactor of the resonance, thus increasing the probability of interactionbetween the wave and the binding agent.

The circulation of light within highly symmetric microscopic structuresoften involves such shape resonances. For dielectric spheres 10-100 μmin size, a particular class of resonances occurs known as whisperinggallery modes. The term stems from similarities with the phenomenon ofcircumferential guiding of faint sounds along the walls of the galleryof St. Paul's Cathedral in London. Bioanalytical and spectroscopicapplications can take advantage of the characteristic of strong surfacelocalization and high quality factors of whispering gallery modes indielectric microspheres and liquid droplets. However, the whisperinggallery modes gradually lose their surface localization properties asthe microsphere size decreases, generally rendering whispering gallerymodes ineffective in a microsphere environment.

For submicron sizes (i.e., less than 1 μm in diameter), one way tomaintain light confinement is to coat the sphere with a surface plasmon(SP) supporting metal film. One characteristic of such a microspherecoated with a metal film is that at certain diameters the total internalreflection angles associated with cavity modes may coincide with the SPRAngle for the metal film, thus resulting in a more efficient form of SPexcitation on geometrically symmetric surfaces. This feature eliminatesthe need for the polarized light source, optical alignment andmechanical scanning found in prior sensors, and allows relaxation of thestringent geometric conditions imposed on planar sensors.

The present invention comprises a novel sensor that may be optimallyused in combination with micro-fluidic systems. Measurements of(bio)chemical concentrations and kinetics of reactions inside a confinedspace such as a micro-fluidic device are very difficult. The presentinvention contemplates a submicron dielectric bead covered with a metalwhich supports surface plasmons, e.g. Au, Ag, Cu. This SPR shows astrong enhancement in transmission of certain wavelengths due to theperiodic boundary conditions created by the geometry of the sensorelement coupled with surface plasmons induced in the metal shell. Thisinventive sensor is sensitive to small changes of the refractive indexof the material at the very surface of the sensor (i.e., within about300 nm) and is much more sensitive than prior far-field sensors anddetection techniques.

Thus, the present invention contemplates a micro-cavity device thatutilizes surface plasmon resonance enhanced by geometric or shaperesonances. For the purposes of the present disclosure, this device willbe referred to herein as a Micro-cavity Surface Plasmon Resonance (MSPR)sensor. In the following description, a spherical cavity resonator hasbeen selected, but it is understood that other symmetric geometricshapes may be used that are capable of sustaining boundary conditionsfor the stationary plasmon resonance wave to travel across the activesurface.

Thus, in one aspect of the invention, the MSPR replaces the propagatingplasmon wave associated with traditional SPR sensors with a stationarywave that travels across the active surface of the sensor element. Inorder to achieve this near-field coupling the dielectric cavityresonator is coated with an SPR-supporting metal of a particularthickness. This metal layer, together with the refractive index of thecavity resonator material, establishes a resonant frequency (orfrequencies) for the cavity resonator sensor element. The dimension ofthe sensor element is then determined in relation to this resonantfrequency. In particular, in one aspect, the sensor element is sized atabout the wavelength of the resonant frequency.

In accordance with the invention, the sensor element or bead is mountedon a light transmissive substrate, such as glass. The substrate and thebead are coated with an SPR-supporting material, such as gold. In afurther feature of the invention, a pinhole is defined at the interfacebetween the bead and the substrate which is free of the coatingmaterial. The size of this pinhole is also calibrated to the resonantwavelength for the sensor, so that the pinhole diameter is less thanthat wavelength. The MSPR sensor further includes a light sourcedirected at the sensor bead that is operable to induce the SPR response.Thus, the light source provides light at the resonant wavelength for thesensor, and may be preferably be monochromatic at the desiredwavelength. The light may be directed at the coated surface of the beador at the pinhole, with a detector positioned to receive lighttransmitted through the MSPR sensor bead.

Due to its small size the MSPR sensors of the present invention can beincorporated into micro-fluidic devices in order to get informationabout the (bio)chemistry occurring inside the micro-fluidic channel.These devices will allow manufacture of compact, disposable sensorswhich can rapidly detect and quantify multiple (bio)chemicals, virusesand bacteria, as well as their concentrations, using small samplevolumes. Thus, the MSPR sensor of the present invention will haveimportant applications in medical diagnostics and therapeutics(especially the diagnosis and treatment of sepsis), in laboratoryinstrumentation for monitoring chemical reactions and in detection ofbiochemical and biological hazards (e.g. bioterrorism or pollution).

In general the MSPR sensor of the present invention can be applied toapplications in which interaction with (bio)chemicals changes therefractive index of the bulk media in contact with the surface of thesensor. In the case of a functionalized detector, the present inventioncan be used in applications in which the chemical interaction causeschanges in thickness or compactness of the self-assembled monolayer thatcovers the surface of the sensor bead and can chemically interact withthe analytes or ligands. Some general (not limiting) applications of theMSPR sensor of the present invention include:

-   -   1. A method to functionalize the detectors inside micro-fluidics        devices.    -   2. Applications detecting molecular species interactions inside        micro-fluidic channels.    -   3. Applications detecting small molecular species.    -   4. Determination of specific binding between molecules.    -   5. Measurements of affinity constants and dissociation constants        of specific molecular pairs, e.g., ligand-receptor pairs,        ligand-antibody pairs.    -   6. Determination of chemical concentrations of analytes inside a        micro-fluidic device.    -   7. Determination of diffusion coefficients of chemicals in        restricted geometries.    -   8. Detection and quantization of molecular species in bodily        fluids, such as blood plasma and urine, in real time.    -   9. Detection and quantization of (bio)chemical or biological        hazards in air and water in real time.    -   10. Detection of molecular species to control release of        therapeutic agents in real time, for instance to control disease        states.    -   11. Detection of hazardous waste or industrial chemicals in air        or water in real time.    -   12. Real time detection of viruses in blood plasma and other        body fluids.    -   13. Determination of blood chemistry in human and veterinary        applications.    -   14. Detection of explosives or explosives/firearms residue.    -   15. Detection of DNA and/or RNA, or detect binding or DNA/RNA        with certain proteins on the order of single cells or at most a        few cells.    -   16. Process analysis and/or control for chemical or biochemical        industrial processes.

One benefit of the present invention is the elimination of thecomplicated optics required for conventional planar sensors. Forinstance, the MSPR sensor of the present invention can use diffuse lightfrom a low-cost light source. The light need not be polarized, filteredor directed. The present invention eliminates the need for fragile, yetbulky, optical alignment components, such as the goniometric mounts inthe prior systems.

A further benefit of the MSPR sensor of this invention resides in itscapability to be integrated into a small package, or chip. The inventiveMSPR sensor allows the light source and the light detector to bepositioned very near the sensor bead array, thereby significantlyreducing the profile of the present MSPR sensor over prior planarsensors.

It is one object of the invention to provide a micro-sensor that iscapable of detecting the presence of target analytes, ligands ormolecules in a fluid. A further object is to enhance the sensitivity andspeed of detection of the micro-sensor.

Yet another object of the present invention is to provide a sensor thatmay provide high throughput detection in micro-environments. Otherobjects and benefits of the invention will become apparent from thefollowing description.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

For the purposes of promoting an understanding of the principles of theinvention, reference will now be made to the embodiments illustrated inthe drawings and described in the following written specification. It isunderstood that no limitation to the scope of the invention is therebyintended. It is further understood that the present invention includesany alterations and modifications to the illustrated embodiments andincludes further applications of the principles of the invention aswould normally occur to one skilled in the art to which this inventionpertains.

In accordance with one embodiment of the invention, a resonantmicrocavity sensor (MSPR) comprises a spherical dielectric microparticle10 supported on a substrate 12, as depicted in FIG. 2. The microparticleis coated with a layer 14 of SPR-supporting material, such as gold, thatis excited through a near-field pinhole 16 defined between themicroparticle and the substrate. The light scattered from the coatedmicroparticle exhibits strong spectral resonances associated with thecoupling of surface-plasmon modes. These resonances can be used forsensing purposes, like the surface-plasmon resonances used for studiesof molecular binding on planar surface-plasmon sensors, but with theadvantages of a submicron footprint and the high quality factors ofmicrospherical resonators, yielding a 100-fold improvement over priorsensors that require optical alignment.

These significant improvements over prior planar sensors areaccomplished, in part, because the sensor of the present inventionrelies upon light transmission rather than reflection. It is known thatreflected light in nano-contexts yields near-field evanescent-wave lighton the far side of the surface of the reflective substrate. The pinhole16 at the interface between the microparticle 10 and the substrate 12has a diameter less than the wavelength of the light directed to thesurface of the substrate, so only near-field evanescent-wave light willpass through the pinhole. However, the light passing through the pinholeis, by itself, insufficient for a sensor to function. Thus, inaccordance with the present invention, the addition of the sphericalresonant cavity above the pinhole converts this near-field light tofar-field light that can be readily sensed or observed. Thesymmetrically shaped microparticle over the pinhole allows thetransmission of light through the pinhole into the resonant cavity toproduce easily observed light transmission above the microparticle. In aspecific embodiment, a laser diode provides light at a wavelength of 590nm, so the pinhole 16 has a diameter less than the wavelength, and morepreferably a diameter of less than 300 nm. In certain examples describedherein, the pinhole diameter established at the contact between the SPRbead and the glass substrate is in the range of 150-200 nm for adielectric micro-particle with a diameter of 771 nm. It is contemplatedthat smaller pinhole diameters will be generated for smaller dielectricmicro-particle diameters.

As expressed above, the MSPR sensor of the present invention does notrequire the complicated optics associated with prior SPR devices thatrely upon surface plasmon waves propagating along a flat substratesurface. In particular, the MSPR sensor shown in FIG. 2 does not requirea light source on a goniometric mount or collimation and filteringoptics for evaluating changes in the SPR angle associated with prior artdevices like the device depicted in FIG. 1. Instead, the MSPR sensor ofthe present invention may be illuminated by light transmittedsubstantially perpendicular to the substrate 12 into the MSPR sensorbeads. Moreover, contrary to the prior art devices of FIG. 1, the lightsource may be situated on either side of the substrate, as explained inmore detail herein.

Furthermore, this freedom from the optical constraints of the priordevices allows the MSPR sensor of the present invention to utilize awide range of light sources at a wide range of frequencies. For thepurposes of the present disclosure, reference to “light” is not limitedto visible light wavelengths. Thus, the light source (or more broadlythe energy source) may provide light in the ultraviolet, visible andinfrared spectral ranges. Although wavelengths outside the UV and IRranges are not presently known to be used in surface plasmon sensors,the invention does not exclude any later discovered energy wavelengthsthat observe the plasmon resonance characteristics of the presentinvention.

Example 1 Fabrication of an MSPR Bead Sensor

The following is a description of one method for laboratory fabricationof the MSPR sensor shown in FIG. 2. It is understood that otherfabrication techniques are possible for specific applications. It isfurther understood that the immediately following description isprincipally for a sensor adapted for research use, rather than forcommercial application, although the same principles may be applied toproduce a commercially viable sensor.

Microscope cover glasses No. 1, 30×24 mm, 156 μm thick, are scored witha diamond and broken into four equal pieces. Also, microscope slides, 25mm×75 mm are scored and broken into two equal pieces. The slides andcover glasses (items 12, 62 and 64, respectively, in FIGS. 2 and 4) arecleaned using a modified version of the well-known RCA cleaning protocol(H₂O₂:H₂O:NH₄OH—2:1:2, warmed to 70° C.) followed by rinsing in DI waterand drying with N₂. The cleaned cover glasses are placed in a dryatmosphere in a bell jar that can be connected to a mechanical pump inorder to create low vacuum inside. Diluted solutions of polystyrenemicrospheres, about 10⁴ beads/μL with diameters 360 nm, 480 nm and 770nm, are prepared in advance and 50 μl of each solution is dispensed oneach piece of cover glass. Due to the cleaning solution, the surface ofthe glass turns hydrophilic. After 2-3 hours of exposure to the vacuumin the bell jar (˜1 torr) the liquid dries out and the beads remainfixed on the cover glass, forming a random, mono-dispersed layer ofbeads. The concentration is chosen so the average distance betweenneighboring beads is sufficiently large (at least 20-50 μm) to avoidoptical cross talk. These samples are sputter coated with a 150 nm layerof gold by exposing them for eight minutes to argon plasma.

An electron-microscope image of a 771 nm polystyrene bead, sputtercoated with 150 nm gold on a glass substrate is shown in FIG. 3. It isunderstood that the sputter coating is capable of producing the pinholeinterface between the bead 10 and the glass substrate 12—in other words,the pinhole is substantially free of the coating material. The lightemitted by the MSPR sensors of the present invention when illuminatedwith white light from underneath the bead sensors is about 100 timesmore intense than the light transmitted through a flat gold layer of thesame thickness.

Example 2 Fabrication of a Micro-Fluidics Chip with MSPR Sensors

According to another embodiment of the invention, a process is providedfor the fabrication of the sensors of the present invention in amicro-fluidic chip, such as the chip 50 shown in FIG. 4. Variations ofthe same protocol will allow fabrication of more complex sensors. Inthis process, the MSPR sensors are mounted within a housing, which inthe preferred embodiments is in the form of a micro-chip. The micro-chipformat for the MSPR sensor allows the sensor to be readily integratedinto micro-systems, such as a micro-fluidics chip described herein.

In accordance with this embodiment, the fluidics chip is made usingphotolithographic technology and chip replica molding inpolydimethylsiloxane (PDMS). The fluidics devices are fabricated usingthe negative-tone photoresist SU-8 as a master to cast PDMS channelstructures. The master substrates are 50 mm×50 mm glass slides. Thesubstrates are cleaned in HCl:HNO3 (3:1), rinsed with de-ionized water,dried with N₂, sonicated in methanol and acetone (2:1), and again driedwith N₂. The master is made with one SU-8 2070 photoresist layer about100 μm thick. The photoresist is spin coated on the glass substrate at3000 rpm for 30 sec and ramped at 120 rpm/sec. After pre-baking on a hotplate for 15 minutes at 65° C. and 90 minutes at 95° C., the photoresistis then exposed to UV light of 365 nm wavelength. The UV exposure systemis equipped with a high pressure Hg arc lamp filtered to pass 360±45 nm,and the exposure dose is 300 mJ/cm². The exposed photoresist ispost-baked on the same hot plate for ten minutes at 65° C. and 30minutes at 95° C. and cooled to room temperature. The master is thendeveloped for ten minutes, rinsed with 2-propanol, and dried with N₂.

The fluidic pattern is transferred to the photoresist through aphotomask drawn using AutoCAD2004 LT and printed on a transparency. Thefluidic pattern in the illustrated embodiment represents a rectangularfluidic chamber 54 (15 mm×10 mm) having two identical channels, an inputchannel 56 and an output channel 58 (5 mm wide and 10 mm long). Thefluidic chamber depth is limited by the depth-of-field of the 60×immersion oil microscope objective used to analyze the sensors. Thefluidic chamber has to accommodate the substrate 12 (156 μm thick in thepresent example) holding the beads 10 covered with the gold layer 14shown in FIG. 2. To provide a fluidic chamber having a depth of about300 μm, the fluidic chamber part of the master is modified by binding apiece of glass substrate 62 identical to that holding the beads.Preferably, the substrates 12 and 62 have substantially the same opticalproperties and thickness.

The silicon elastomer kit contains a polymer base and curing agent thatare mixed in a 10:1 ratio for five minutes. A tape barrier is placedaround the mold to hold the elastomer mixture, and the elastomer ispoured onto the master. The PDMS in the mold is placed under low vacuum(˜1 torr) for one hour to enhance fluidic pattern replication and curedby heating at 120° C. for twenty minutes. The PDMS substrate is thenseparated from the master, and access holes for fluid connections to thechannels are punched through the elastomer with a 16 G needle.

At the bottom of the fluidic chamber of the PDMS chip 50 the substrate12 holding the beads covered with gold is attached to the ceiling of thefluidic chamber 54 of the PDMS chip 50 with a drop (50 μL) of PDMS. Thesubstrate is placed with the sensors facing away from the PDMS mold andexposed to the inside of the fluidics chamber. The binding is finalafter ten minutes baking at 90° C.

The fabricated PDMS substrate and a 25 mm×50 mm No. 1 cover glass 62 arethen permanently joined after being exposed to air plasma for 40 secondsprior to contact. To increase the rigidity of the chip 50 and toeliminate mechanical perturbations in the flow, a half microscope slide64 (25 mm×38 mm) is permanently bound on top of the chip using the sameair plasma technique. In this example, the depth of the fluidic chamberis estimated to be less then 50 μm in one specific embodiment so thatthe sensors can be brought into the focus of a 60× oil immersionobjective with a working distance of 200 μm.

Example 3 Operation of the Micro-Fluidics MSPR Sensor Chip

In one method of using the micro-fluidics chip 50 (FIG. 4), fluidconnections from the fluidics chip to fluid reservoirs, such as asyringe or a fluid pump, are made using 1.6 mm OD polypropylene tubing.The flows are controlled by adjusting the height of the reservoirconnected to the input channel 56 relative to the height of thereservoirs connected to the output channel 58, or controlled by thefluid pump, so that a stable flow of about 1 μL/s is achieved. After thechip is connected to the reservoirs it is placed on a piezo-driven stagecapable of motion in all three directions (3D) that can position thesensor chip in space with a precision of 10 nm. The whole ensemble isplaced under an inverted microscope and microscope objectives of 40× and60× are used to collect and analyze the signal coming from a singlesensor.

The functionality and sensitivity of the MSPR sensor 50 may be evaluatedusing an experimental set-up shown in FIG. 5. In one experiment, thesensitivity of the sensor to vapors is tested. The substrate holdingsensors is placed on a 3D-piezo-driven stage that can position thesensor in space with a precision of 10 nm. The whole assembly is placedon the stage of an inverted microscope and microscope objectives of 40×are used to collect and analyze the signal coming from a single sensor.The light coming from the sensor is fed through a parallel port into amonochromator driven by a data acquisition interface unit. Spectra inthe visible range from 400 nm to 800 nm may be recorded on a PC with aresolution of 2 nm and 1 sec detector integration time.

In accordance with one embodiment, the experimental set-up includes atube connected to a bubbler placed in the proximity of the sensor and N₂is purged through a solution of water:200 proof ethanol (2:1). Thevapors are periodically turned on and off in order to check the sensor'sresponse to the stimulus. Spectra of the light emitted by the dry sensorand the wet sensor are recorded, as shown in FIG. 6 a. The peak mostsensitive to vapor concentrations is preferably chosen for recording thetime-series of the transmitted light, as shown in FIG. 6 b. (Theabscissa in both graphs corresponds to the ratio of light intensitybetween the SPR bead and the flat film surrounding the bead). The graphin FIG. 6 a shows the spectral shifts in the light transmitted through a771 nm Au-coated bead due to water (the continuous line corresponding to50% humidity at ambient atmosphere) and ethanol vapor adsorption (thedotted line corresponding to 75-80% humidity with the vapor accessopen). The graph in FIG. 6 b shows the measured sensor response(wavelength=715 nm) to cyclic humidity changes between 50% and 80%. Thearrows represent the instant when the vapor access was opened (down) orclosed (up).

To further verify that the sensor is indeed sensitive to surfacemodifications, alkanethiol adsorption from ethanol may be employed as aprobe. Formation of a single monolayer is known to occur in ˜100 minutesat 10 mM concentrations. A comparison of the spectra is provided in FIG.7 a together with the spectral transmission of the flat gold film forthe same conditions. The shift of the spectral transmission through aflat film is less than the signal that spherical cavity SPR sensor isexpected to record. The spherical cavity sensor of the present inventionis thus more sensitive than the flat film sensors of the prior art. Thespectral shifts in FIG. 7 a persist after the dodecanethiol solution isflushed with pure ethanol, indicating that irreversible adsorption ofalkanethiol has occurred. The shifts are thus due to the formation of amonolayer at the gold surface. Upon measurement of adsorption kineticsand fitting with a first order exponential decay (FIG. 7 b), a timeconstant is found for the film formation of 382±7 s at a 100 mMdodecanethiol concentration. Note that while the signal in FIG. 7 bcorresponds to a single monolayer about 1.5 nm thick, thesignal-to-noise ratio is good enough to detect binding of fractions of amonolayer.

Example 4 Alternative MSPR Sensor Configuration

In Example 1 described above, the sensor responds to excitation throughthe pinhole 16 (FIG. 2). In an alternative embodiment, the sensor isconfigured for excitation through the head of the sensor array, asdepicted in FIG. 8. In this embodiment, cover glasses No. 1, 24 mm×50 mmand 160 μm thick are used as a substrate. NIST standard polystyrene780±5 nm diameter beads were prepared in concentrations of about 10⁴beads/μl in methanol. (Methanol was chosen in this example because ithas a very low superficial tension coefficient relative to water andtherefore produces a suitable randomly mono-dispersed array of beads).In the example, 70 μl of beads solution was dispensed on each coverglass, providing a bead density of about 5000 beads/mm². After beingdried at 1 torr vacuum for an hour, the substrates were sputter coatedwith a 140-150 nm layer of gold in a masked region of about 10 mm longand 3 mm wide. The substrates were burnt in air plasma for about 3minutes to ensure that the sensor was clean. Both substrates and PDMSmolds were exposed for 45 seconds to air plasma prior to contact.

In this embodiment, the MSPR sensor is positioned within amicro-fluidics structure that permits fluid flow across the random arrayof beads, as reflected in FIG. 8. The structure may be configured as aT-shape with two micro-fluidics channels that are 50 μm wide and 20 μmdeep connected to a common channel 100 μm wide and 20 μm deep. The MSPRsensors are disposed within the common channel. This micro-fluidicsstructure is molded into the PDMS elastomer and holes are formed in theelastomer to access the two flow channels. The flow channels areconnected to two corresponding reservoirs placed at different heights.In this example, flow through the channels is thus accomplished simplyby hydrostatic pressure and is on the order of 100 μm/sec. Of course, inother embodiments or commercial versions, fluid flow through themicro-fluidics structure may be accomplished in any manner, such as by afluid pump.

In this example, a number of MSPR sensors are mounted on the floor ofthe common channel of the micro-fluidics device, again as shown in FIG.8. Rather than illuminate the sensors through the pinholes (as in theprevious example), the sensors are illuminated through the head of thedevice—i.e., through the spherical surface of the beads. It was foundthat the sensor of this example exhibited a resonant response similar tothat in the example depicted in FIG. 2, except that the embodiment ofFIG. 8 experienced a greater signal-to-noise ratio.

One benefit of the embodiment of FIG. 8 is that enclosure of themicro-fluidics chip is facilitated. In order to enclose themicro-fluidics chips of the present invention, both the glass substrateand the PDMS mold are exposed to air plasma which modifies the chemicalstructure to bond the two media. Since the MSPR sensors are very smalland the spacing between the beads is in the range of 10-100 μm, applyingthe gold layer, such as by sputter coating, is problematic. Inparticular, it is difficult to apply the gold layer to the beads onlyand not to the glass substrate. On the other hand, gold does not bondwell to the glass substrate. The embodiment of the present exampleallows a continuous compact layer of gold to be coated onto the sensorbeads and the bottom glass substrate of the sensor. A layer of amaterial having an affinity for both glass and gold may be added to theglass substrate. In a specific embodiment, the material may be chromiumapplied at a thickness of about 1-5 nm. Alternatively, the substrate maybe subject to a chemical treatment to improve the adherence between thegold and the substrate. The PDMS may then be applied and flows wellthrough the microchannels between the sensors. Post-baking the PDMSmolds at 80-100° C. overnight cures the polymer and eliminates anyvolatiles or loose polymer chains that might infiltrate the gold layersputtered on the glass substrate.

Example 5 Functionalization of MSPR Sensors

Covalent functionalization on the gold surface of the sensor shown inFIGS. 2 and 8 allows the sensor to be covered with different targetanalytes, ligands or molecules, particularly biomolecules of highinterest. For the purposes of the following disclosure, the term“target” or “targets” shall be used to generically refer to the targetanalytes, ligands or molecules that are intended to be detected by thesensor. It is understood that these “targets” may include biomolecules,such as proteins, RNS, DNA and enzymes, as well as elements other thanbiomolecules, such as viruses, bacteria, non-biological chemicals, etc.However, it is understood that these “targets” have the ability to bindwith other molecules provided on the MSPR sensors of the presentinvention and do so in a way that affects the resonant characteristicsof the sensors.

In accordance with certain embodiments of the invention,functionalization of the gold layer is accomplished in this example bytwo different chemistries in the form of a respective monolayercovalently reactive to proteins. The first chemistry isDithiobis(N-succinimidyl propionate) (DSP, DTSP), also known as Lament'sreagent, which is a homobifunctional thiol-cleavable cross-linker thatadsorbs onto gold surfaces through the disulfide group. DPS is a highlyhydrophobic compound that is soluble in dimethylsulfoxide (DMSO) ordimethylformamide (DMF). A water soluble cross-linker sulfo-DPS (DTSSP)is used to avoid interaction between the DMSO or DMF and the goldsurface. The functionalization reaction is illustrated in FIG. 9. Inparticular, the disulfide bond breaks and reacts with the gold surface.

DTSSP is a semi-stable amine-reactive NHS-ester that is proteinreactive. In this example, the reaction is evaluated using two differentworking buffers phosphate buffer pH 5.8 and DI water. The reactionkinetics results in a monolayer molecule of 281.52 Da and about 0.6 nmthick at a time constant of 105±8 sec. and a signal-to-noise ration of5.5

The second chemistry used for the MSPR sensor's functionalizationincludes Carbodiimide coupling reagents. The reaction involved with thischemistry occurs in three steps. The first step is a reaction of a zerocross-linker with the gold surface, as shown in FIG. 10 a. In this firststep, the cross-linker is 3,3′-Dithiodipropionic acid (DTDPA) that has adisulfide bond that easily breaks in the presence of gold. Thiscross-linker ends in a carboxyl group that permits carbodiimidecoupling. The DTDPA reaction kinetics yields a molecule of only 104 Daand a 0.5 nm monolayer at the surface of the sensors. A carbodiimidemediator 1-Ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC) isemployed to readily react with nucleophiles. The EDC solution isprepared in ethanol because the EDC can hydrolyze very quickly. The samesolution also contains an amine-reactive ester, such asN-Hydroxysuccinimide ester (NHS). In the complex reaction illustrated inFIG. 10 b the EDC and NHS promote a carbodiimide coupling reaction thatconverts the carboxylic acid into a reactive intermediate that issusceptible to attack by amines. Thus, the final product is aminereactive and ready to bind proteins to the surface of the MSPR sensor.

The reaction kinetics of this second chemistry was found to form amonolayer of 98.1 Da and about 0.5 nm thick on top of the zerocross-linker with an estimated time constant of 28.5±0.9 sec. at asignal-to-noise ratio of 16.2.

Example 6 Protein Binding to a Functionalized SPR Sensor

One significant application of the sensors of the present invention isas a bio-sensor. Thus, a sensor functionalized in the manner describedin Example 5 may be used to detect certain protein molecules that arecapable of binding to the functionalized chemistries. Two importantbio-molecules are glucose oxidase (Gox) and glucose (Glu).

Gox is a very large molecule made of two identical subunits having atotal MW of 160,000 Da. Thus, Gox provides a good test to assess theability of the MSPR sensor and micro-fluidics device to respond tobinding of large molecules. Gox is known to bind to gold in a particularorientation and to an EDC/NHS activated gold surface in a differentorientation. In this example once the gold surface of the MSPR sensor isactivated with the amine-reactive NHS-ester groups, as described above,reaction to proteins is simple but the reaction time constants willdepend upon the size of the protein. For a DTSSP functionalized sensor,the reaction time constant was found to be 562±35 sec with a signal tonoise ratio of 3.85. This reaction covered the sensor surface with amonolayer of about 10 nm thickness. In this reaction it was determinedthat the MSPR sensors of the present invention exhibited a sensitivityof 42 zepto-moles/SPR sensor, or expressed as Gox mass covering thesensor a detectability of 6.7 femtograms/MSPR sensor. Variations of thisprocedure were implemented to monitor the Gox activity under theinfluence of a flow of β-D+Glucose 100 mM in PBS 1×, or a flow of Glu 1mM in PBS 1× (to simulate normal glucose concentration in human blood),or a flow of L-Glucose, or a flow of 2-Deoxy-D-Glucose (2-DxGlu). Thedevice in the present example was able to detect the enzymatic activityof Gox in the presence of β-D+Glucose 100 mM and 1 mM (except that forthe latter case the response was much slower), but no enzymatic activityresponse was recorded for Gox exposed to L-Glucose or 2-DxGlu.

The above examples demonstrate the efficacy of the MSPR sensor andmicro-fluidics features of the present invention in detecting large andsmall targets, including bio-molecules such as important proteins. Inparticular, the MSPR sensors of the present invention can be configuredto a footprint of less than 1 μm and are still capable of detectingspecific binding of zeptomoles of unlabeled targets. In accordance withthe present invention, the light source in the optic setup may be alaser diode. In the examples, the selected laser diodes resonated at awavelength of 590 nm; however, it is contemplated that other small laserdiodes may be used at other wavelengths. It is believed that a laserdiode resonance at a wavelength of 632.8 nm may help optimizeperformance of the SPR sensors of the present invention.

It is contemplated that light sources other than the above-describedlaser diode may be used. For instance, in certain alternativeembodiments, a light source may incorporate an optical filter operableto limit the transmitted light to a desired wavelength(s). The opticalfilter may be tuned at the time of installation of the MSPR sensor to aspecific resonant frequency. Alternatively, the optical filter may bepositioned at the detector side of the sensor.

The selection of optical detectors can enhance functionality andefficiency of the MSPR sensors of the present invention. In one specificembodiment, the detector may be a low dark current silicon avalanchephotodiode (APD) photon counting detectors. Alternatively, for detectingmultiple targets in parallel, a CCD camera or other pixel orienteddevice may be used. The detectors and associated electronics candetermine a baseline resonant peak for the MSPR sensors to calibrate thesensor. In use, the detectors may determine whether the resonant peakhas shifted (red or blue), which is a direct indication that the targethas bound to the resonant surface of the MSPR sensors.

The invention contemplates detectors that are qualitative—i.e., thatsimply detect the presence of a particular target—or quantitative—i.e.,that detect the level or change in level of the target. In the lattercase, a quantitative analysis can be particularly valuable to measurethe change in analyte concentration over time. For instance, changes incertain toxins in a patient's blood may be monitored, rather than simplydiscrete instantaneous level, thereby facilitating early diagnosis of aharmful medical condition. The example herein regarding detection ofsepsis may benefit from this quantitative approach. Similarly, at homequantitative monitoring of blood sugar levels may be used for earlierdetection of diabetic conditions.

In the embodiments described above, the gold layer is sputter coatedonto the MSPR beads and the glass substrate. Since adhesion between goldand glass is poor, the manufacturing process may include sputtering athin layer of chromium onto the glass before adding the gold layer,since chromium binds well to glass and gold binds well to chromium. Forhigh throughput manufacturing, both layers may be applied by a twin headsputter coater to avoid the need to break the vacuum around thesubstrate.

In the above examples, fluid flow through the micro-fluidics device wasaccomplished by hydrostatic pressure only. Alternatively, themicro-fluidics sensor chip may incorporate micro-valves and peristalticpumps to control fluid flow and sample delivery. The use of thismicro-fluidics technology will also allow the micro-fluidics sensors ofthe present invention to process small sample volumes, on the order of 2μl. Thus, a micro-fluidics MSPR sensor in one embodiment of theinvention may be configured as shown in FIG. 11. The sensor 70 includesa MSPR substrate 72 with a T-shaped micro-fluidics structure 74 mountedthereon. The T-shaped structure 74 operates in the manner describedabove to direct fluid from the channels 75 a, 75 b of the structure tothe common channel 76 over the MSPR sensors. A second level of themicro-sensor 70 includes the fluid control components. In particular, amicro-fluidic pump 78 is provided at the discharge end of the commonchannel 76. In specific embodiments, the pump may be peristaltic,thermal, or piezo-actuated. Each channel 75 a, 75 b is provided with acorresponding micro-valve 79 a, 79 b to control fluid flow through therespective channel into the common channel 76. In a single analytedetection sensor, such as the micro-sensor 70 shown in FIG. 11, onechannel 75 a and valve 79 a controls flow of the sample into the commonchannel, while the other channel 75 b and valve 79 b controls flow ofthe functionalization solution.

It is contemplated that the micro-fluidics components may beelectronically controlled to operate in a pre-determined sequence forfunctionalizing the MSPR sensor array and analyzing a fluid sample. Inparticular, valve 79 a may be closed and valve 79 b opened to permitintroduction through channel 75 b of functionalization solutions, suchas the functionalization composition as described above. Once the SPRsensors are functionalized, a buffering solution may be introducedthrough the channel 75 b. The valve 79 b may then be closed and valve 79a opened to accept the sample fluid through channel 75 a to contact thefully functionalized SPR sensor array. Of course, it is contemplatedthat the functionalization step may occur remote from the sampleanalysis—i.e., in the preparation of a pre-packaged biologicalmicro-sensor.

In addition, the pump and micro-valves may be controlled as necessary toensure sufficient formation of the monolayer of the target on thefunctionalized MSPR sensor. For instance, in the Gox example above, theformation of a 160,000 Da monolayer about 10 nm thick was detected witha time constant of about 562 seconds. Thus, the flow of test fluidthrough the micro-fluidics chamber must be adequate to ensure theformation of a significant and detectable monolayer of the target.

Another aspect of the fluidics element of the inventive sensors isdependent upon the nature of the fluid sample being evaluated. Inparticular, a complex sample requires cleaning and pre-concentrationbefore analysis to ensure accurate detection results. Such complexsamples include human blood, which may be evaluated for certain proteinsas described in the functionalization examples above, and natural water,such as water from a river being evaluated for the presence of dangerouspathogens. Pre-cleaning and pre-concentrating a biological sample mayoccur prior to introduction into the MSPR sensor system. For instance,centrifugation may be used to clean a fluid sample, but centrifugemachines are not adapted for a micro-fluidics environment. Large-scalesample testing, such as a drinking water purity monitor, may be amenableto this scale of pre-cleaning and pre-concentrating. However, onefeature of the present invention is that is very well suited formicro-fluidics applications in which the entire sensor and associatedsample fluidics are present on a single small chip.

Thus, the present invention contemplates the addition of micro-fluidicfiltration and pre-concentration modules that are integrated onto theMSPR sensor chip. Thus, a system 80 shown in FIG. 12 may incorporate amicro-fluidic filter module 82 and a pre-concentration module 84upstream of the MSPR sensor chip, such as the chip 70 illustrated inFIG. 11. In this embodiment, the upstream modules are connected to thefluid sample channel 75 a and valve 79 a.

The micro-fluidic filter 82 in a specific embodiment includes a porousmembrane sandwiched between opposing PDMS molds. The flow area of thefilter depends upon the fluid sample being tested. For instance, afilter area of about 2.5 cm² is sufficient for low volume filtering,such as up to 1 ml of blood. Larger filter areas may be required forhigher volume, or higher flow rate sampling.

In general, filtration removes some of the targets that are desired tobe detected. For instance, many proteins will non-specifically bind tofilter membranes. Thus, in some cases a pre-concentration module 84 maybe interposed between the filter module 82 and the sample channel 75 aof the micro-fluidics sensor chip. A variety of pre-concentrationapproaches may be acceptable, such as electrophoresis, capillaryseparation, functionalized magnetic bead, isotachophoresis, columnseparation or photo-activated polycarbonate (PPC) micro-fluidics chips.

The small size and the accuracy of the MSPR sensor chip of the presentinvention allows the fabrication of sensors with throughput andmassively parallel processing capabilities that greatly exceed thecapabilities of current sensors and biosensors. In particular, the MSPRsensors of the present invention can be configured to detect thousandsand even millions of targets, all on a single small sensor chip. Asshown in FIG. 13 the sensor chip includes a plurality of MSPR beads on asingle chip that may be arranged in randomly mono-dispersed arrays or inregular arrays. The arrays of MSPR sensors may be produced usingphotolithography and/or holographic optical tweezing, or any othersuitable technique for placing microscopically small objects onto aglass substrate. However, one feature of the present invention is thatmillions of the micro-sized MSPR beads may be completely randomlydispersed on the substrate using currently available technology. Asexplained below, in spite of this random dispersion of MSPR beads,sensors made according to this embodiment of the invention may be fullyfunctionalized to detect a vast number of targets.

In order to accommodate the need to detect multiple target, currentplanar SPR sensor technology requires uniformly distributed SPR elementsto ensure adequate detection capabilities for multiple targets. Therelatively low sensitivity of these current sensors dictates that asufficient number of SPR elements be associated with predetermined“spots” in which all elements are functionalized to a particular target.However, the ability to accurately place uniformly distributed SPRelements is very limited, generally not exceeding a 100 by 100 grid ofelements. This limitation, coupled with the accuracy limitations of thecurrent planar sensors, ultimately limits the number of discrete targetsthat can be detected to less than about 1000, which ultimately severelylimits the range of applications for these sensors. For instance, genetherapy and human genome mapping projects yield millions of targets fordetection. Using the current planar technology, hundreds of the bulkysensors would be necessary for projects of this nature.

On the other hand, the capability exists to randomly disperse themicro-beads utilized in the sensor of the present invention. However,until the present invention, there has been no way to capitalize on thisability to populate a sensor substrate with millions of SPR elements,each capable of being functionalized individually or in groups of spots.In accordance with the present invention, one method of achieving thisdiscrete functionalization is to operate on groups of sensors by flowingreagents over specific bands of the sensor chip using micro-fluidics. Inother words, as seen in FIG. 13 the chip 72 may be divided into multiplebands, such as the four lengthwise bands 86 a-d. A micro-fluidics systemmay then flow a specific reagent along each band to commonlyfunctionalize each MSPR sensor along the band. This approach limits thedegree of functionalization to the number of bands on the chip overwhich the various reagents may accurately flow. In one specificembodiment, the MSPR chip may be divided into about twenty bands, eachwith different functionalization so that a like number of targets may beearmarked for detection.

In another approach, individual MSPR sensors may be precisely selectedfor specific functionalization. One manner of achieving this individualfunctionalization may be by use of a photo-activation boundcross-linker, such as photo-biotin. However, this method is inherentlyslow since only a few SPR sensors may be functionalized at a time.Another more versatile approach is to use a micro-spotter for makingmicro-arrays of SPR bead sensors, in a manner similar to prior ink jetprinters. Some micro-spotter printers are capable of placing ink dropsto a resolution of 600×600 dpi, with dot sizes in the range of 30 μm at45 μm spacing and a volume of only 10 μl. Even more accurate ink jetprinters are capable of resolutions of 4800×4800 dpi with each ink dothaving a diameter of only 5 μm. This printing technique may be adaptedto functionalize selected MSPR sensors or groups of sensors, resultingin functionalized spots, such as the spots 88 shown in FIG. 13. Eachspot may pertain to a different target.

In yet another approach, discrete multiple target functionalization maybe achieved using a multi-pin spotter. This multi-pin spotter mayprecisely apply the cross-linker or reagent directly to and only on theMSPR beads. The specifically functionalized beads may be in clusters orrandomly dispersed throughout the entire field of MSPR beads.

In a further approach to functionalization that is well suited tomassively parallel processing, the MSPR beads may be functionalizedusing a mask. The mask limits the application of the cross-linker orreagent to the MSPR beads disposed within spots 88 on the substrate. Itis contemplated that the functionalized spots will encompass randomnumbers of the randomly dispersed MSPR beads on the array over an areathat is significantly larger than the beads themselves. Thus, in aspecific embodiment, the functionalized spots may occupy an area about30 μm in diameter, whereas the MSPR beads have a diameter of about 770nm. A micro-spotter capable of dispensing reagents in quantities as lowas 10 μl may be used to functionalize the beads in each spot. The sensorchip may include a bar code 86 or some other readable signatureidentifying the various functionalizations as well as spotscorresponding to each functionalization. As described below, the barcode 86 may also contain calibration information corresponding to theresponsive signals generated by the detector 90 (FIG. 14).

With the sensor construction as thus far described, a plurality ofrandomly dispersed MSPR beads populate the substrate, with collectionsof beads commonly functionalized to form spots 88. In the specificexample shown in FIG. 13, eighteen such spots are depicted; however, itis contemplated that hundreds, thousands and even millions of such spotsmay be defined on a given sensor chip. An operational sensor chiprequires a light source and some form of detector to sense the resonantresponse at each spot. Thus, in accordance with one embodiment of theinvention, a stack forming the micro-sensor may appear as shown in FIG.14 with the MSPR sensor chip 72 sandwiched between a detector 90, whichmay be a CCD array, and a light source 96, which may be an LED. It isunderstood that various optical conditioning elements may be integratedwith the light source and/or detector, such as an optical filter toimprove signal/noise ratio. The optical conditioning element may alsoinclude a wavelength filter or different discrete wavelength filterscorresponding to specific spots 88 or individual MSPR beads.

In accordance with one feature, the detector or CCD array may be mappedinto a grid 92, with each pixel 93 of grid containing a CCD capable ofsensing light transmission through the MSPR sensor chip 72 andconfigured to generate a signal indicative of that light transmissionfor subsequent processing. This mapped grid 92 overlays the sensor chip,as shown in FIG. 14, or alternatively the spots 88 may be regarded asprojected onto the mapped grid, as illustrated in FIG. 13. Optimally,the detector grid is fine enough so that each spot 88 may be projectedonto multiple pixels 93 of the grid. It is expected that each pixel mayoverlay several MSPR beads, although the number of beads correspondingto each pixel will vary due to the random distribution of the beads onthe substrate.

Calibration of the detector proceeds first by identifying an optimumpixel or pixels reading transmission data from each spot 88. Thus, in aspecific example, a particular spot may fully encompass four pixels 93and partially encompass five additional pixels. The MSPR chip isilluminated by the light source 96 and the measured intensity at each ofthe pixels corresponding to the spot is evaluated. The pixel registeringthe greatest response is selected as the pixel corresponding to thespecific spot, which in turn corresponds to a specificfunctionalization. That selected pixel will likely map onto the largestnumber of MSPR beads relative to the other pixels, hence its greaterresponse relative to the other pixels. The output from the CCD withinthis selected pixel may then be calibrated in relation to the intensityand/or wavelength of the light source 96. This same process is repeatedfor all of the other functionalized spots 88. Thus, in the specificexample, for the eighteen functionalized spots (FIG. 13), eighteenpixels 93 on the mapped grid 92 of the detector 90 may be identified sothat the calibrated output of each pixel will be evaluated. Thiscalibrated output may be written onto an on-board memory or transmittedto a peripheral memory device and/or processor. A calibration table withthe calibration data for each of the mapped pixels may be maintained ina memory and accessed by the peripheral processor. The bar code 86 maythus provide an identifier for extracting the proper calibration tablefrom multiple tables stored in memory. The calibration table mayidentify which pixels to read from the detector and how to interpret theoutput signal from each pixel. The peripheral device applying thecalibration data may be configured to obtain the necessary data from aglobal database, such as through an Internet link.

It is contemplated that additional pixels may also be associated with aparticular spot, with appropriate modifications to the calibration ofthe corresponding output responses. It should be appreciated that insome cases the output response for a given pixel may result from lighttransmission through only one MSPR bead present within a given spot andaligned with a given pixel, while for another pixel the lighttransmission may be measured through several MSPR beads. The randomdistribution of beads means that the number of MSPR beads used togenerate an output signal corresponding to each functionalized spot isalso random. However, the calibration step described above can ensurethat the targets can be quickly and accurately detected. The highsensitivity of each MSPR bead in the MSPR sensor of the presentinvention means that even a single MSPR bead may be sufficient for aparticular functionalized spot and detector array pixel.

It can be appreciated that the device illustrated in FIG. 14 may openrealms of target detection unavailable with prior sensor devices. Asexplained above, a single MSPR sensor chip may be functionalized tothousands of targets in a small package. The small size of the sensorsof the present invention allows the formation of massively parallelarrays of sensors for DNA, RNA and protein detection. The use ofmicro-fluidics with the sensor chip allows for a continuous flow of testfluid across the sensor chip 70. This micro-fluidics feature facilitatesthe massively parallel sensor arrays and provides an avenue forreal-time accurate sensing of chemical and biochemical conditions.

A particularly beneficial usage is in real-time detection of targets inthe blood stream. One important application of the multi-channelembodiments of the present invention is in the detection of sepsis.Sepsis is a major source of mortality in post-surgery recovery and intrauma victims. Treatment of sepsis is largely limited to antibioticsand palliative measures to support heart, lung and kidney function.According to data collected in 2001, sepsis syndrome affects anestimated 751,000 patients in the United States each year, of whom383,000 (51.1%) received intensive care. Mortality has been estimated at215,000 deaths nationwide, increasing with age from 10% in children to38.4% in those 85 years and older. The cost per case averages about$22,000, which means almost $17 billion annually. Early detection ofsepsis and rapid intervention (within two to four hours of onset)greatly reduces mortality and debilitation in survivors. However, nocurrent method exists to monitor patients for the onset of sepsis. Inmany cases the medication produced for sepsis treatment failed due tothe lack of instrumentation capable to continuously monitor cytokineslevels in patients' blood.

Sepsis syndrome is the body's systemic inflammatory response toinfectious stimuli. Endotoxins—such as lipopolysaccharide (LPS) fromGram-negative bacteria, peptidoglycans and flagellan from Gram-negativeand Gram-positive bacteria, lipotechoic acid from Gram-positivebacteria, mannan from fungi, and other antigens from infectiousagents—stimulate macrophages and monocytes to release tumor necrosisfactor alpha (TNF-α), followed by a cascade of cytokine release. Duringthe first period of sepsis (especially the first eight hours), excessiveinflammatory response can cause massive organ damage, especially tokidneys and heart, but also reaching the liver, lungs and brain,requiring artificial support of blood pressure and ventilation. Thisorgan damage often causes debility or mortality months or years afterthe acute phase of sepsis.

The release of pro-inflammatory mediators was originally thought to belargely uncontrolled. However, subsequent investigations havedemonstrated that TNF-α also stimulates leukocytes to releaseanti-inflammatory cytokines, including IL-10, IL-1 and transforminggrowth factor-beta (TGF-β), which inhibit the synthesis ofpro-inflammatory cytokines and exert direct anti-inflammatory effects onmonocytes, macrophages, and endothelial cells. This compensatoryanti-inflammatory response syndrome (CARS) is intended to localize whatwould otherwise be an uncontrolled pro-inflammatory response to theinfection throughout the body. Unfortunately, the anti-inflammatoryresponse often surpasses the pro-inflammatory response in the laterphases of sepsis, resulting in immunoparalysis—i.e., the inability tomount an effective immune response to additional infectious insults.

Thus, a minimally-invasive device, which could be attached to allpostoperative and post-trauma patients which could monitor the onset andprogress of sepsis, allowing for much earlier and more directedintervention, and ultimately reducing both mortality and debility insurvivors. In accordance with a particular embodiment, a micro-fluidicsdevice 100, shown in FIG. 15, is provided that can be used tosimultaneously analyze a set of six chemicals that play an importantrole in sepsis. (However, the micro-fluidic structure can be modified toaccommodate the analysis of more or fewer chemicals at the same time, ormay be modified for different driven flows, flow velocities or analysisset-ups). One manner of diagnosis of the onset of sepsis and itsprogress involves monitoring the chemicals, TNF-α, IL-1, IL-6, IL-10,IL-13 and TGF-β, at the same time and in real time. Thus, as shown inFIG. 15, the micro-fluidics sensor 100 includes a set of six channels102-107 (each 50 μm wide, 100 μm apart and 11 mm long in a specificembodiment) that come together into a micro-fluidic chamber 110 (2 mmlong and 600 μm wide), with each channel corresponding to a particularmonitored chemical. The spacing is chosen so that diffusion of moleculesof about 20 kDa will not interfere with molecules in the neighboringchannels. These six channels are used to functionalize the sensor array120 with specific anti-bodies for the above-mentioned chemicals. Theflow of different antibodies through the micro-fluidic chamber willcreate a series of six parallel antibody stripes 112, functionalizingthe chip. The chip 100 in this specific embodiment is designed for aflow of the antibody at a minimum 400 μm/s. This treatment and specificcalibration can be done prior to usage of this chip to evaluate asample.

A transverse channel 116 of 200 μm wide intersects the micro-fluidicchamber 110. This channel is wider to permit more viscous and fastercoagulating fluids, like blood plasma, to pass though. The region ofintersection covers an area of 200 μm×600 μm. The transverse channel 116includes an inlet 117 through which the blood is introduced, and a wasteoutlet 118. Similarly, the micro-fluidic chamber 110 includes a wasteoutlet 111.

This sensor array 120 is preferably oriented at the intersection of thechamber 110 with the transverse channel 116 and preferably aligned withthe channel. In one specific embodiment, the sensor 120 may include alinear array of six SPR bead sensors 124 or sensor spots, with eachindividual sensor or sensor spot corresponding to a particularfunctionalization and aligned with the corresponding antibody strip 112.

The micro-fluidic device 100 depicted in FIG. 14 may be fabricated usinga poly-dimethylsiloxane (PDMS) substrate and a cover glass in the mannerset forth below. The devices are fabricated using negative-tonephotoresist SU-8 as a master to cast PDMS channel structures. The mastersubstrates are 50 mm×50 mm glass slides. The substrates are cleaned inHCl:HNO₃ (3:1), rinsed with nanopure water, dried with nitrogen,sonicated in methanol and acetone (1:1), and dried with nitrogen again.The master is created with two SU-8 photoresist layers. A firstunder-layer (an 18-20 μm thick layer of SU-8 2010) is used to promoteadhesion of the channel structure to the substrate and a second thickerlayer (60-80 μm thick layer of SU-8 2070) of photoresist is used tocreate the channel structure. Both layers are processed identicallyexcept that the first layer is exposed without a photomask. Thephotoresist is spin coated on the substrate at 1000 rpm for 30 secondsand ramped at 40 rpm/second. After prebaking on a hot plate for oneminute at 65° C. and two minutes at 95° C., the photoresist is thenexposed to UV light. The proposed channel design is transferred to thephotoresist through a photomask drawn using AutoCAD 2004 LT and printedon a transparency using a high resolution laser printer at 8000 dpi. TheUV exposure system is equipped with a high-pressure Hg arc lamp filteredto pass 360±23 nm, and the exposure dose is 300 mJ/cm₂. The exposedphotoresist is postbaked on the same hot plate for one minute at 65° C.and three minutes at 95° C. The master is then developed for fiveminutes, rinsed with 2-propanol, and dried with nitrogen.

The silicone elastomer kit contains a polymer base and curing agent thatare mixed in a 10:1 ratio for five minutes. A tape barrier is placedaround the mold to hold the elastomer mixture, and the elastomer ispoured onto the master. The PDMS on the mold is placed under low vacuum(˜1 torr) for one hour to enhance channel replication and cured byheating at 120° C. for twenty minutes. The PDMS substrate is thenseparated from the master, and access holes for fluid connections to thechannels are punched through the elastomer with a 16 G needle.

At the bottom of the PDMS mold, across the micro-fluidic chamber 110 atthe intersection with transversal channel 116, the linear array of MSPRbeads 120 may be produced using photolithography and/or holographicoptical tweezing, or any other suitable technique for placingmicroscopically small objects onto a glass substrate like using amicromanipulator and a laser tweezers system. The micromanipulator isloaded with a solution of 10³-10² beads/μL, precision-size-standardbeads. This concentration is chosen so that the MSPR beads of the array120 are dispensed having a spacing of about 50-100 μm. An optical lasertweezers can be used to hold the bead in place until the liquid driesand the next bead will be dispensed with the micromanipulator and heldwith the tweezers until the liquid dries and so on and so forth. Oncethe beads 124 are placed, the PDMS substrate is sputter coated through awindow of 1 mm×1 mm placed above the intersection. The beads are coveredwith 150 nm of gold. The PDMS substrate and glass cover glass are thenpermanently joined after being exposed to air plasma for 40 secondsprior to contacting.

In one embodiment, a sepsis detector device may include acatheterization tube connected intravenously to the patient and to apumping system to periodically draw a small volume of blood into thesensor device 100. The blood passes through a disposable filter toextract the plasma and the plasma is supplied to the disposable sensor100 through a channel 116. The output from the sensor provides a readingof the cytokine concentrations in the plasma. The waste blood passesthrough channel 118 to be collected in a disposable biohazard-labeleddiscard tube. The small size of the MSPR sensor and sensor chip of thepresent invention allows the sensor device 100 and catheterization tubeto be in place as long as the patient is under medical care. Control ofthe peristaltic pump to draw blood into the sensor chip may beelectronically controlled to occur at pre-determined intervals or inresponse to some other medical condition sensor. The detector of thesensor may be coupled to the same controller to generate an alarm if theparticular agents are detected.

Elevated levels of IL-10, IL-13 and TGF-β indicate incipient sepsis,while elevated levels of IL-10, IL-13 and TGF-β indicateimmunoparalysis. If the sensing time for each of the sensors is toolarge, a single channel may not be able to detect the cytokine levels inreal time. In that case an array of channel detectors may be fabricatedon a single disposable chip and the blood supply switched betweenchannels at five minute intervals. A controllable microvalve may be usedto alternately supply blood and clean sterile medium to the sensors toreset them for the next measurement.

Classes of agonist and antagonist drugs have been developed to controlthe various cytokines involved in sepsis. However, none of these drugsare widely used because physicians have no way to monitor their effects,which vary greatly from patient to patient and over time. The result isthat the current treatment of choice includes inflammatory suppressorsand enhancers that are given in either insufficient or excessive doses,both of which may be lethal to the patient. The MSPR sensors of thepresent invention would allow physicians to supply a tailored cocktailof agonists and antagonists which would suppress immune response earlyin infection and enhance it in late infection, while maintaining thecytokines at optimal levels at all times.

The same principles for detecting sepsis conditions may be applied tothe interactive detection of other medical conditions, as well as aninterface to collateral therapeutic devices. With appropriatefunctionalization, a single or multiple-sensor device may be used tomonitor patient status during extended treatments. For instance, a MSPRsensor chip and micro-fluidics system in accordance with the embodimentsdescribed herein may be incorporated into a dialysis system, or otherdevice that continuously draws blood or other fluids from a patient fortreatment. The MSPR sensor chip may be integrated into a continuousblood monitoring system to detect targets in real-time that areindicative of oncoming problems, such as heart attack, stroke, kidneyfailure and the like.

A second embodiment of the multiple sensor array devices may be providedthat comprises of a set of syringe tubes containing cytokine regulatorydrugs and controlled by the output of the cytokine detector prescribedabove. This device would ultimately supply a controlled dosage ofmultiple cytokine regulators to the patent via an intravenous (IV) drip,continuously changing the supply of agonists and antagonists to keep thepatient's cytokines at optimum levels. In the above blood monitoringexample, the real-time detection of targets indicative of the onset of aheart attack, for instance, may be used to provide immediate real-timedosing in response to the onset of that condition.

This same interventional treatment may be employed to stave off sepsiswhen detected as described above. In this instance, certain anti-sepsistreatments rely upon the action of a particular protein to inhibit thecreation of certain target molecules. However, the treatment itself maybe immune-suppressant, so the treatment must be carefully administered.Real-time detection of target levels by the MSPR and micro-fluidicsensors of the present invention allow prompt and accurateadministration of the anti-sepsis treatment. A similar approach may beimplemented to reduce the toxicity of chemo-therapy or HIV treatments,or for other treatments that create target blood-borne markersindicative of the onset or presence of unwanted side effects.

The disposable MSPR sensor devices described above are suitable for manyother applications. For instance, the present invention may be adaptedfor public and private drinking water testing. The MSPR sensors may befunctionalized to detect various organic and inorganic contaminants,toxins, cellular organisms and viruses. The sensors may be positionedwithin the water supply to continuously monitor the water flow for theselected targets. Since the devices of the present invention rely uponlight detection devices, such as the CCD array 90 described above, anelectrical signal is generated that may be evaluated and used toinitiate a predetermined response, such as a sensible alert.

Devices based on these sensors can be developed for detection ofbiohazards, noxious chemicals, neurotoxins, explosives, or HIV or otherviruses or bacteria in blood, plasma or other body fluids. The presentinvention allows the sensors to be small enough to be portable andeasily disposable. In the illustrated embodiments, the sensor chip fitswithin a 50 mm×50 mm area. Specific apparatus can be adapted for airportor homeland security use or for use in water treatment plants orfactories. Depending on the use, automated systems for connecting to theinput reservoirs of the chip can be included and additional chemicalscan be analyzed at the same time and on the same chip.

The MSPR sensors and micro-fluidics of the present invention may also beadapted to monitor chemical reactions or bioreactions. The micro-chipsof the present invention may be integrated into fluid flow lines ordirectly within chemical reactors or bioreactors to detect certaintarget products of the reactions or to detect the chemical conditionswithin the reactors that may impact the reaction. The MSPR sensors maybe used to optimize the reaction conditions or determine when thereaction is complete. This specific embodiment may have beneficialapplication as part of process control for drug or chemical fabrication,especially to control the purity of the resultant product.

Micro-fluidics devices endowed with the sub-micron cavitysurface-plasmon biosensors of the present invention overcome severaldeficiencies in prior sensing techniques and devices. Combining theproperties of the micro-fluidics devices with the sensitivity of theMSPR bead sensor extends the boundaries of the lab-on-a-chip ideal byincreasing detection abilities inside the micro-fluidics chip orconfined spaces.

Current devices monitor molecular interactions and molecular kineticsusing planar SPR or the older ELISA kits. In order to excite surfaceplasmons on a planar metal surface certain restrictions must be obeyed.In particular, the source of light must be p-polarized and a precisecritical angle of incidence must be obtained in order to produce amaximum coupling between incoming photons and surface polaritons. Thesensor of the present invention combines the sensitivity of surfaceplasmons with the resonant properties of a spherical sensor. Besides aboost in sensitivity, the invention relaxes constraints on the geometryand polarization of the light source. Moreover, the sensor has afootprint of a square micron or less, which makes it well-suited forminiaturization (having an active area of about one thousand timessmaller than the present state-of-the-art SPR planar sensors) whileincreasing sensitivity and improving the ability to integrate intomicro-fluidic structures. Furthermore, the MSPR sensors of the presentinvention work in transmission compared with the prior SPR sensors thatwork in reflection. Due to this difference, a sensing micro-fluidicschip incorporating the MSPR sensor of this invention can be placed veryclose between the light source and the sensing window of the detector,resulting in a very compact, robust and inexpensive handheld device.

Micro-fluidics devices are currently the most sophisticated technologyfor dealing with small quantities of analyte (ranging from picoliters tomicroliters), and for very precise control of flows and gradients. Theyare appropriate for multiple replica molding and new configurations ofchannels and set ups can be produced at very low cost, making themsuitable for single-use devices. This property makes them convenient inmany areas of research, especially for medical applications. They aresuitable for massively parallel processing of chemicals, which can savea huge amount of time, especially in analyzing very complex samples, forexample, but not limited to, blood plasma, body fluids, toxic waste,foods, etc. The only time constraint is the reaction time between thespecific molecular species in the sample and the receptors bound to thesurface of the functionalized MSPR detector.

The MSPR sensor of the present invention takes advantage of theseproperties of micro-fluidics devices. It has better sensitivity thanprior devices due to the coupling of the surface-plasmon with thegeometry of the sensor. While current SPR sensors can only detect largemolecules, the SPR sensors of the present invention can detect large andsmall molecules with good sensitivity. The smallness of the sensorsallows control, detection and analysis to be achieved on a single chipand to take advantage of the ability of micro-fluidics devices toconduct multiple parallel analyses, which can shorten the analysis time.Moreover, these devices can be disposable and can be produced cheaply.

In the illustrated embodiments, the microspheres are coated with gold.It is contemplated that the spheres may be coated with other metals,such as silver, copper, or gold alloys. However, current experimentationsuggests that spectral resonances in the transmitted light occur onlyfor gold-coated spheres or beads, which is believed to be due to thesurface plasmon coupling. The film thickness of the gold coating may beadjusted depending upon the application of the particular MSPR sensor.However, it has been found that increasing film thickness causesblue-shifting of the observed resonances, at least for the symmetric(low frequency) mode. Conversely, it has also been found that thefrequency for the high frequency anti-symmetric mode is red-shifted withincreases in film thickness. Moreover, the spherical geometryimplemented in the present invention preferentially excites thesymmetric SP modes, thereby minimizing red-shift effects.Experimentation further suggests that some peaks, such as the peak at623 nm for a 770 nm bead coated with a 150 nm gold layer, exhibits muchless sensitivity to the metal film thickness than other peaks.

While the invention has been illustrated and described in detail in thedrawings and foregoing description, the same should be considered asillustrative and not restrictive in character. It is understood thatonly the preferred embodiments have been presented and that all changes,modifications and further applications that come within the spirit ofthe invention are desired to be protected.

In the illustrated embodiments, the micro-particles forming the MSPRsensor are spherical in shape to form a spherical resonant cavity.However, other symmetric geometric shapes may be utilized for the beadshape. For example, the bead may have an elliptical shape or bemulti-faceted like a dodecahedron, provided that the shape can sustainperiodic boundary conditions for the stationary plasmon wave to travelacross the surface of the bead.

The various materials and dimensions set forth for the illustratedexamples may also be modified while still maintaining the functionalityachieved by the MSPR sensor and micro-fluidics systems of the presentinvention. Modifications to the materials and dimensions of the MSPRsensor must still fulfill the primary object of the MSPR sensors of thepresent invention, namely to detect targets Moreover, the modificationsmust not interfere with the shape or geometric resonant characteristicsthat are used to enhance the SPR resonance features of the micro-cavitysensor. The detection capability of the MSPR sensors of the inventionrelies upon binding the target to a coupling reagent layer that ititself bound to the SPR-supporting coating, and ultimately upon thechange in optical response.

It is believed that for most targets and coupling reagents thewavelength of the applied light is not critical. On the other hand, theSPR-supporting layer is, by definition, wavelength dependent since thesurface plasmon resonance occurs in that layer. Thus, it is believedthat modifications to the materials and dimensions of the MSPR sensorare centered on the selection of the SPR-supporting layer material andits characteristic wavelength. In the illustrated embodiments, thatmaterial is gold which has a wavelength of 510 nm. In accordance withcertain aspect of the invention, this wavelength determines the diameterof the micro-particles or beads 10 and the pinhole 16. The bead diameteris also a function of the refractive index of the dielectric material.

In alternative configurations, the SPR-supporting coating material maybe silver, copper, or other non-gold SPR-supporting material, withappropriate changes in coating thickness. Since silver and copper eachhave a different SPR characteristic wavelength, the selection of eithermetal as the material for coating 14 will result in a change in diameterfor the micro-particle 10 and the pinhole 16. In accordance with certainembodiments of the invention, the micro-particle diameter would be sizedto about the characteristic wavelength of the silver or copper coating,while the pinhole diameter would be fixed at less than that wavelength.Similarly, the coating thickness may be modified with a commensuratechange in the micro-particle and pinhole diameters.

To the extent that the MSPR sensor dictates the characteristicwavelength, the light source and transmitted light detector (such as thesource 96 and detector 90 in FIG. 14) may be selected accordingly. Incertain embodiments, white light may be acceptable, while in otherembodiments it may be desirable to select a monochromatic light sourcecentered at the characteristic wavelength of the MSPR sensor. Preferablythe light detector is calibrated to the characteristic wavelength.

With respect to material selection for the MSPR sensors andmicro-fluidics chips of the present invention, the materials in theabove examples and embodiments are illustrative. While the MSPR beadsare described as formed of polystyrene, other light transmissivematerials may be used, such as glass or aluminum oxide. The selectedmaterial is most preferably dielectric and has an index of refractionsimilar to polystyrene. Of course, as indicated above the index ofrefraction of the bead material affects the optical response andresonant mode of the MSPR, along with the SPR-supporting coating.

Likewise, the material forming the housing or chip around the MSPR beadsand substrate may be different from the PDMS material identified in theillustrated examples and embodiments. Preferably, the material issubstantially light transparent and exerts only a minimal influence onthe optical and resonance characteristics of the MSPR sensor.

With respect to applications or uses of the MSPR sensors andmicro-fluidic sensors of the present invention, the foregoing examplesand embodiments are not intended to be limiting. It should beappreciated that the present invention permits the rapid and accuratedetection of a wide range of targets, whether in small sample volumes orin continuous flow systems. The present invention also permitssimultaneous detection of hundred, thousands and even millions oftargets in a single micro-sensor or in a massively parallel array ofsensors. Thus, even as the present invention may greatly enhance currentdetection techniques, it will likely lead to new techniques and analysesnot yet contemplated.

1. A sensor chip for mounting between a light source and a detector fordetecting the presence of a target analyte, ligand or molecule in a testfluid, comprising: a light transmissive substrate; a functionalizedsurface plasmon resonant (SPR) element mounted on said substrate, saidSPR element functionalized to bind with the target, where the SPRelement includes a light transmissive bead formed in a geometric shapethat can sustain periodic boundary conditions for a stationary plasmonresonance wave to travel across the outer surface thereof, and a pinholeis defined at the interface between said SPR element and the substrate;a housing formed of a light transmissive material and defining a fluidcavity for flowing the test fluid over said SPR element, said housingfurther defining at least one fluid inlet and at least one fluid outletin communication with said cavity; and at least one micro-fluidiccomponent associated with said housing and operable to control fluidflow through said cavity.
 2. The sensor chip of claim 1, wherein saidmicro-fluidic component is a micro-fluidic pump.
 3. The sensor chip ofclaim 2 wherein the micro-fluidic pump is mounted within said housing.4. The sensor chip of claim 1, wherein at least one of said fluid inletsincludes a micro-fluidic valve.
 5. The sensor chip of claim 4 whereinthe micro-fluidic valve is mounted within the housing for controllingfluid flow through said cavity.
 6. The sensor chip of claim 1, whereinsaid housing further defines at least one additional fluid inlet incommunication with said cavity.
 7. The sensor chip of claim 6 whereinsaid additional fluid inlet includes a micro-fluidic valve.
 8. Thesensor chip of claim 1, wherein said micro-fluidic component includes afilter.
 9. The sensor chip of claim 1, wherein said micro-fluidiccomponent includes a pre-concentration module.
 10. The sensor chip ofclaim 9 wherein the pre-concentration module is a microfluidicpre-concentration module.